Chronically implantable devices may provoke inflammation and/or fibrosis from tissue trauma or tissue response to the foreign body. See, e.g., Reichert et al., Handbook of Biomaterial Evaluation, Ch. 28 Biosensors, pp. 439-460, (Von Recum A., editor) (1999); Wisniewski et al., J Anal Chem 2000; 366 (6-7) (p. 611-621).
Implanted devices may also cause other unwanted bioreactions. For example, recently, researchers have stated that drug-coated stents might cause adverse reactions leading to blood clot formation in some patients. See Lagerqvist et al., Long-Term Outcomes with Drug-Eluting Stents versus Bare-Metal Stents in Sweden, New England Jnl. Of Medicine, Mar. 8, 2007.
Other chronically implantable devices that may invoke unwanted bioreactions are biosensors. For example, in order to maintain near normal blood glucose levels (70-120 mg/dL), diabetic patients widely use over-the-counter glucose meters, which require finger pricking to obtain blood samples several times a day. The pain (Lee et al., 2005), inconvenience, and discomfort of self-monitoring of blood glucose (SMBG) are frequently obstacles to effective patient compliance and optimal management of diabetes. During the past 20 years many kinds of continuous glucose monitoring systems have been studied including sensors implanted in the subcutaneous tissue (Moussy et al., 1993; Johnson et al., 1992; Koudelka et al., 1991; Bindra et al., 1991; Pickup et al., 1989; Shichiri et al., 1986; and Ertefai et al., 1989), sensors implanted in the vascular bed (Armour et al., 1990; Frost et al., 2002), and determining glucose concentration in interstitial fluid sampled using a micro dialysis device (Ash et al., 1992; Meyerhoff et al., 1992; Moscone et al., 1992). Although several studies of implantable glucose sensors have been reported, it is believed that none of these biosensors are reliably capable of continuously monitoring glucose levels during long-term implantation. Progressive loss of sensor function occurs due in part to biofouling and to the consequences of a foreign body response such as inflammation, fibrosis, and loss of vasculature (Reichert et al., 1992; Reichert et al., 1999; Sharkawy et al., 2007). Some researchers have modified the surface of the sensors to reduce membrane biofouling in vivo. In an approach to reduce protein adsorption, Quinn et al., 1995 used poly(ethylene glycol) (PEG) in a polyhydroxyethylmethacrylate (PHEMA) matrix. Since the PEG chains tend to stand up perpendicular to the membrane surface, they provide a water-rich phase that resists binding of many protein molecules. Rigby et al., 1995 and Reddy et al., 1997 reduced protein adsorption by using diamond-like carbon, so-called “inert” materials. Shichiri et al., 1988 incorporated an alginate/polylysine gel layer at the sensor. Shaw et al., 1991 reported improvement in biocompatibility of a biosensor coated with PHEMA/PU (polyurethane). Wilkins et al., 1995 and Moussy et al. introduced NAFION (perfluorosulphonic acid) membrane (Du Pont), to reduce “biofouling” on the surface of the sensor and reduce interference from urate and ascorbate (Moussy et al., 1993; Moussy et al., 1994a; Moussy et al., 1994b; Moussy et al., 1994c). Armour et al., 1990 coated their sensor tips with cross-linked albumin and Kerner et al., 1993 developed cellulose-coated sensors to improve sensor blood compatibility. However, it is believed that none of these approaches has been satisfactory for long term, stable glucose monitoring.
Collagen and its derived matrices are used extensively as natural polymers in the biomedical field including tissue engineering due to its low antigenicity, its biodegradability and its good mechanical, haemostatic and cell-binding properties (Sheu et al., 2001; Pieper et al., 2002; Chvapil et al., 1973; Pachence et al., 1996; and Lee et al., 2001). In order to devise strategies for using collagen in the development of advanced biomaterials for biomedical engineering, it is typically desired to confer mechanical strength and resistance to enzymatic (collagenase) degradation resistance with chemical or physical cross-linking strategies. There are several strategies for cross-linking collagen-based biomaterials. Glutaraldehyde (GA) is the most widely used as a cross-linking agent for collagen-based biomaterials (Sheu et al., 2001; Barbani et al., 1995). However, GA and its reaction products are associated with cytotoxicity in vivo, due to the presence of cross-linking byproducts and the release of GA-linked collagen peptides during enzymatic degradation (Huang-Lee et al., 1990; van Luyn et al., 1992).
In order to avoid in vivo cytotoxicity and subsequent calcification of GA cross-linked collagen, several alternative compounds have been examined as potential collagen cross-linking agents (Khor et al., 1997; Sung et al., 1996) such as polyepoxy, hexamethylene diisocyanate (HMDI), 1-ethyl-3-(3-dimethylamino-propyl)carbodiimide (EDC), and ultra-violet (UV) or gamma-ray irradiation. Koob et al. recently described a process for cross-linking of type I collagen fibers with nordihydroguaiaretic acid (NDGA), a plant compound with antioxidant properties (Koob et al., 2002a; Koob et al., 2002b; Koob et al., 2001a; and Koob et al., 2001b). Koob et al. showed that NDGA significantly improved the mechanical properties of synthetic collagen fibers. In addition, they showed that NDGA cross-linked collagen fibers did not elicit a foreign body response nor did they stimulate an immune reaction during six weeks in vivo.
The extent of cross-linking and choice of cross-linking agent may also affect the porosity and pore size of the scaffold and may influence fibrous capsule thickness, blood vessel density, and the location of vessels within the three-dimensional porous scaffold (Joseph et al., 2004). Large pore scaffolds (greater than 60 micron pore size) allow deep penetration of capillaries and supporting extracellular matrix (ECM). Sharkawy et al., 1997 showed that after four weeks of subcutaneous implantation in rat, a well-organized collagen matrix typical of a foreign-body response encapsulated non-porous implants, while the porous polyvinyl alcohol (PVA) implants produced less fibrous and vascularized tissue capsules.